Source filter for radiographic imaging

ABSTRACT

A system for analyzing biological structures by photon absorptiometry is disclosed, including a radiation source emitting photons and at least one source filter which operates to reduce the emission of high energy photons, thereby creating a sharp drop-off at the high end of the detected energy spectrum. The filter radiation is received by a detector which determines the spatial intensity of the radiation over the image area and preferably has a characteristics response which is relatively insensitive to low energy photons, such that the source filter and detector cooperate to measure the intensity of photons within a narrowed band of energy levels.

BACKGROUND OF THE INVENTION

The technical field of this invention is radiology and, in particular,bone absorptiometry by radiographic measurements.

The depletion of bone mineral content, typically referred to asosteoporosis, is a common consequence of a variety of diseases andnatural aging process. In addition to metabolic bone diseases and aging,bone minerals can be lost as the result of drugs, stress, dietarydeficiencies, pregnancy or lactation. When skeletal bone mass dropsbelow the level necessary to provide mechanical support, the depletionof bone mineral content becomes an important cause of morbidity,particularly in elderly patients.

Unfortunately, at present there are no reliable and inexpensive systemsfor gauging bone mineral content (BMC) with any high degree ofprecision, particularly during the early stages of osteoporosis or othermineral depletion disorders when dietary supplements and therapeuticagents may reverse the course of demineralization and preventdebilitating fractures or otherwise slow the progress of the disease.

Conventional methods for determining bone mineral content typicallyinvolve measurements of radiation absorption in the bone. U.S. Pat. No.3,715,588 issued to Rose on Feb. 6, 1973, is illustrative of a prior art"bone scanner" in which a collimated X-ray beam is passed through a bone(e.g., the wrist) and detected by a radiation detector mechanicallycoupled to the X-ray source. The system scans back and forth across thebone to produce a complete measurement of the bone and surroundingtissue. Because of inherent differences in tissue and bone absorption,bone density (and, hence, mineral content) can be inferred from alogarithmic ratio of the intensity of radiation detected aftertransmission through the two media.

One problem with such scanning systems is the time required for scanningcan lead to inherent blurring of the image if patient motion occurs.Additionally, the spatial resolution of conventional detectors does notpermit identification of the same bone area in repeated scans. Thesedefects lead to BMC errors on the order of five percent or more forindividual patients, precluding effective use of radiographic imagingtechniques for monitoring Progressive charges in bone mineral content.

There exists a need for more precise bone absorptiometry systems thatcan be used by the general medical community. In particular, systemswhich could operate to irradiate the entire image area simultaneously toform "snap shots" of the bone structure would satisfy a long-felt needin the art by greatly reducing the problem of patient motion. Moreover,higher spatial resolution systems which reduce the measurement errorsassociated with current bone densitometry imaging would also address anunmet need in this field.

SUMMARY OF THE INVENTION

A system for analyzing biological structures by photon absorptiometry isdisclosed, including a radiation source emitting photons and at leastone source filter which operates to reduce the emission of high energyphotons, thereby creating a sharp drop-off at the high end of thedetected energy spectrum. The filter radiation is received by a detectorwhich determines the spatial intensity of the radiation over the imagearea and preferably has a characteristic response which is relativelyinsensitive to low energy photons, such that the source filter anddetector cooperate to measure the intensity of photons within a narrowedband of energy levels.

By presenting the detector with a narrow band of energy levels, higherspatial resolution is achieved. In one illustrated embodiment, theradiation source is a broad band X-ray source emitting X-rays at energylevels up to about αKeV and the source filter creates a sharp drop-offof photons having energy levels above about 50 Kev. The detector can bea multiple wire proportional chamber (MWPC) device filled with anionizable gas, or alternatively, a scintillation screen or aphotosimulatable luminescence measuring device. In one illustratedembodiment, a MWPC device which is insensitive to most photons havingenergies less than 35 KeV.

In another aspect of the invention, the system can also include one ormore additional source filter elements to further narrow the energy bandof the photons. In one embodiment, two or more metal filter elements,ranging from 0.01 to 1.0 millimeters in thickness, are employed tosharpen the energy band and provide for highest transmissivity in therange of about 42 to about 48 KeV.

It has been discovered that a composite K-edge filter using 2 or more,preferably 3, pure metals stacked on one another, is particularly usefulin modifying the raw spectrum of a radiation source to achieve greaterspatial resolution in the detector. It has been found that metals withatomic numbers from 45 to 52 (usually two metals) combined with at leastone metal whose atomic number ranges from 57 to 65 give an optimumspatial resolution in Xe gas detectors. For example, cadmium (Cd),silver (Ag), and gadolinium (Gd) can be particularly useful incombination. Similarly, cadmium (Cd), tin (Sn) and samarium (Sm) formanother useful combination filter.

The result of using the composite filter is twofold: first, by blockinglow energy photons, it reduces the radiation dose otherwise absorbed bythe patient and, second, by narrowing the energy band of the transmittedphotons, it dramatically diminished the projected electron range in thedetector, itself, thereby improving spatial resolution.

Another result of the composite filter lies in its ability to reduce theerror typically inherent in bone mineral content (BMC) measurements dueto the presence of fat in the soft tissues of the wrist and forearm.Bone mineral content measurements typically involve measurement of theintensity of photons through a tissue equivalent bolus which is the samethickness as patient's limb and which contains a material having anabsorption characteristic which is the same as tissue free of fat.However, the fat content can be a significant (e.g., five to fifteenpercent) absorptiometry factor in heavy individuals; and when it varies,it will lead to errors in measurement of BMC changes from year to year.Since the composite filter removes a large percentage of the photons inthe range 20 to 35 Kev where the difference between the absorptioncoefficients for fat and tissue is largest, the observed differences inphoton intensity through fat-containing tissue will be reducedsubstantially for the filtered x-ray spectra.

The invention will next be described in connection with certainillustrated embodiments. However, it should be clear that variouschanges, additions and subtractions can be made without departing fromthe spirit or scope of the invention.

BRIEF DESCRIPTION OF THE DRAWINGS

FIG. 1 is an isometric view of a radiographic imaging system accordingto the present invention;

FIG. 2 is a schematic block diagram of a radiographic imaging systemaccording to the invention;

FIG. 3 is a more detailed schematic diagram of the radiation source anddetection of the system of FIG. 2;

FIG. 4 is a more detailed schematic illustration of the beam collimatorof FIG. 3;

FIG. 5 is an exploded schematic illustration of the beam filter of FIG.3;

FIG. 6 is a schematic side view of a multi-wire detector module usefulin the system of FIG. 2;

FIG. 7 is a more detailed schematic illustration of the cathode andanode grids of the detector of FIG. 6;

FIG. 8 is a graph showing a typical raw energy spectrum of an X-raysource, such as shown in FIG. 3;

FIG. 9 is a graph showing a typical energy transmission spectrum of abeam filter, such as shown in FIG. 3;

FIG. 10 is a graph showing a typical filtered beam spectrum afterpassage through a beam filter, such as shown in FIG. 3;

FIG. 11 is a graph showing typical detection efficiencies versus beamenergy for a detector, such as shown in FIG. 3;

FIG. 12 is a graph showing the overall energy response of a radiographicimaging system according to the invention;

FIG. 13 is a more detailed schematic top view of a limb positioning andcalibration device of the system of FIG. 2;

FIG. 14 is a schematic side view of the positioning and calibrationdevice of FIG. 13 in a closed position;

FIG. 15 is a schematic side view of the Positioning and calibrationdevice of FIG. 13 in a closed position;

FIG. 16 is a schematic block diagram of the data read-out circuitry ofFIG. 6;

FIG. 17 is a more detailed diagram of a gated event digitizer as shownin FIG. 16;

FIG. 18 is a graphic illustration of the priority encoding operationperformed by the circuitry of FIG. 16 to obtain an event coordinate;

FIG. 19 is a schematic block diagram of data acquisition circuitryuseful in the system of FIG. 2;

FIG. 20A is a timing diagram for a programmable logic device as shown inFIG. 19 and further illustrated schematically in FIG. 20B;

FIG. 21 is a schematic illustration of a data flow path within thecircuitry of FIG. 19;

FIG. 22 is a schematic illustration of a patient X-ray image (showingnon-imaged bones in phantom) in accordance with the present invention;

FIG. 23 is a more detailed view of an X-ray image illustrating a methodof identifying bone reference points according to the present invention;and

FIG. 24 is another detailed view of an X-ray image illustrating a methodof defining bone analysis regions according to the present invention.

DETAILED DESCRIPTION

FIG. 1 is an isometric view of a radiographic imaging system 10 foranalyzing biological structures according to the present inventionconsisting of a radiation source 12, which transmits a beam of radiationto detector module 14. Disposed between the radiation source 12 and thedetector module 14 is the limb of a patient 16, which is secured inplace by a limb positioning apparatus 18 and registration means 15A, 15Bwhich prevents the limb from pivoting about the positioning apparatus 18in operation. The detector module 14 and the limb positioning apparatus18 and the registration means 15A, 15B conveniently can be incorporatedinto a housing 20 so that the patient can sit while the procedure isbeing conducted.

FIG. 2 is a more detailed schematic block diagram of the radiographicimaging system as shown, including radiation source 12, detector module14, positioning apparatus 18, data acquisition module 22, dataprocessing module 24, data display module 26 and data storage module 28.Radiation emitted by the radiation source 12 passes through thepatient's limb (e.g., the wrist) 16 and is detected by detector module14 and converted therein to digital data. The data is then transferredto the acquisition module 22 where a radiographic image of the patient'slimb is built up over time. The images formed in the acquisition modulecan be analyzed in the data processing module 24 (e.g., to detectosteoporosis or other changes in bone mineral content), displayed by thedisplay module 26 and/or stored for subsequent analysis or display indata storage module 28.

FIGS. 3-7 are more detailed schematic illustrations of the radiationsource 12 and detector module 14. As shown in FIG. 3, radiation source12 includes a radiation head assembly 30 which houses a radiationemitter 32, such as a radioisotope or, preferably, an X-ray tubeoperable to generate photon radiation having maximum energy levels, forexample, on the order of about 45 to about 70 KeV and, preferably, about55 KeV. Such X-ray tubes are available commercially from a variety ofsources. One such commercially available tube is the Model OIX-15 dentalX-ray tube (Eureka Company, Chicago, Ill.).

Also shown in FIG. 3 is a detector module 14 which preferably is an areadetector capable of determining the spatial intensity of radiationemitted by the radiation source 12 and transmitted through the patient'slimb 16. One such area detector is a multi-wire, proportional chamberdetector, described in more detail below in connection with FIGS. 6 and7. Other area detectors which can be employed in the present inventioninclude scintillation screen devices, employing gadolinium sulfate orlanthanum sulfate crystals and associated electronic imaging (i.e., CCD)elements, or photosimulatable luminescence (PSL) measuring devices.Detectors of these types are commercially available, for example, fromKodak (Rochester, N.Y.) or Fuji (Tokyo, Japan).

The radiation head assembly 30 shown in FIG. 3 also includes a threadedbarrel 33 into which can be inserted a collimator assembly 34 and afilter assembly 38. In FIG. 4, a more detailed schematic illustration ofa collimator assembly 34 is shown consisting of a machined metal (i.e.,lead) body 37 having a central hole 35 and peripheral threads 36 so thatit can be screwed into the threaded barrel 33 of the radiation headassembly.

FIG. 5 is an exploded schematic illustration of the beam filter assembly38, including a filter housing 40 and cover plate 42 which can beconnected by bolts 44 on the housing through holes 46 on the cover platevia nuts or other suitable attachment means. Disposed within the filterassembly is one or more filter elements. In the illustration, three suchfilter elements 48, 50 and 52 are shown. In one preferred embodiment ofthe invention, three filter elements of gadolinium, silver and tin, eachhaving a thickness on the order of about 0.08 millimeters can beemployed.

More generally, in the invention, it is preferable to employ at leastone filter element which operates to reduce the transmission of highenergy photons in conjunction with a detector module having acharacteristic response which is relatively insensitive to low energyphotons, such that the source filter and the detector module cooperateto measure the intensity of photons within a narrowed band of energylevels.

The principal of operation is that the detector is chosen such that ithas an energy response which increases efficiency rapidly above adefined energy level, known as the "K-edge", for a particular detectormaterial. Xenon gas-based detectors have a K-edge at about 35 KeV.Barium-based PSL imaging plates have a similar K-edge at about 37 KeV.Gadolinium scintillation screens have a K-edge at about 50 KeV. Thefilter is chosen, such that its K-edge is higher than that of thedetector, such that the resultant system acts as a bandpass filter witha lower limit defined by the K-edge of the detector and an upper limitdefined by the K-edge of the filter.

In one preferred embodiment, at least one filter element has anelemental composition consisting of one or more of the lanthanidemetals, such as lanthanum, cerium, praseodymium, neodymium, promethium,samarium, europium, gadolinium, terbium, dysprosium, holmium, erbium,thulium, ytterbium, lutetium or alloys thereof. This filter element canbe employed alone or in combination with one or additional filterelements having an elemental composition of silver, cadmium, indium, tinor alloys thereof. As noted above, one particularly useful combinationof filter elements is the ternary combination of gadolinium, silver andtin. Each filter element can have a thickness ranging from about 0.01 toabout 1.0 millimeters, preferably ranging from about 0.05 to about 0.5millimeters.

The collimator 34 is employed to eliminate scattering effects that maybe caused during the filtering processes and produces a beam ofradiation of appropriate size and shape for passage through the patientlimb undergoing radiographic imaging to the detector module.

In FIG. 6, a multiwire proportional chamber (MWPC) device 80 isillustrated as part of detector module 14. The chamber 80 is filled witha gas 82, which typically comprises xenon (Xe) and one or more buffer oradditive gases, and contains an X-cathode grid 84, a Y-cathode grid 86and an anode grid 88. The MWPC device 80 can be incorporated into ahousing 78 having a window or otherwise radiation-transparent uppersurface 79. Voltage source 90 produces a positive charge on the anodegrid 88. The cathode grids 84, 86 are connected to data readoutcircuitry 92 to sense ionization events which occur when the photonradiation ionizes the xenon gas molecules within the chamber 80 and whenthe electrons, resulting from such ionization, drift to the positiveanodes where charge multiplication takes place.

In FIG. 7, the operation of the MWPC device 80 is further illustratedschematically. In the simplified illustration, five X-cathode and fiveY-cathode lines 84, 86 are disposed in orthogonal relationship (each ofwhich is actually composed of a series of parallel wires which areganged together) forming a 5×5 coordinate matrix. The X-grid 84 and theY-grid 86 are separated from each other by an anode grid 88. When aphoton of sufficient energy strikes a xenon gas molecule, the gasmolecule is ionized, producing an electron-ion pair. The electron driftstowards the anode where the intense-field around the anode results incharge multiplication. The resulting avalanche 94 of electrons is sensedas a positive charge on the nearest wires of both the X-cathode and theY-cathode grids, as well as a negative charge on the anode grid. Thecoincidence of the X-cathode signals, X_(i), the Y-cathode signals,Y_(i), and the anode signal, a, permits the detector to accuratelypinpoint the location of the radiation-induced ionization event.

In practice, the grids are larger and finer than shown in FIG. 7 toprovide higher spatial resolution. For example, one embodiment for wristimaging can employ 160 parallel Y-cathode wires and 80 parallelX-cathode wires, which are ganged as pairs to provide a 40×80 locationgrid. The spacing between adjacent wires can be approximately 1millimeter. Alternatively, the cathode grids can be constructed onradiation-transparent (e.g., polyimide) printed circuit boards, forexample, having copper strips 4-5 millimeters wide separated from eachother by gaps of about 0.1 to about 0.5 millimeters in width.

The anode grid 88 can be similarly constructed of a plurality of thinparallel (e.g., 20 micrometer gold plated tungsten) wires. The anodewires, however, can be all wired together and, in operation, are all atthe same voltage determined by the voltage source 90. The voltage willdepend upon the pressure of the Xe gas within the chamber, as well asthe spacing between the anode wires, and the anode-to-cathode spacing.For a typical pressure of 3 atmospheres and a spacing between anodewires of 2 millimeters, a positive voltage on the order of about 2,500to about 5,000 volts relative to the cathode wires can be employed.

FIG. 8-12 illustrate how the filter elements and detector cooperate tonarrow the energy band of detected radiation in the present invention.For radiographic imaging with the illustrated system, the preferredenergy band ranges from about 35 KeV to about 50 KeV. (The graphsillustrate X-ray energies from about 27 KeV to 55 KeV; photons havingenergies below 27 KeV are typically absorbed by the patient's body anddo not reach the detector.)

When an X-ray tube or the like is used as a radiation source, photonshaving a spectrum of energy levels are produced as illustrated in FIG.8. The upper limit on the spectrum (i.e., 55 KeV in FIG. 8) is obtainedby controlling the voltage applied to the X-ray tube. It is preferableto narrow this raw spectrum so as to achieve higher spatial resolution.

In the illustrated multi-wire proportional chamber (MWPC) detectionsystem, Xe gas is used as the radiation detecting element. Photons ofenergy below 34 Kev can ionize the Xe atoms and produce L-shellelectrons, having energies equal to the energy of the photon minus 5.0Kev (L-edge of Xe); likewise, photons with energy greater than 34.5 Kevcan ionize the Xe atoms and produce K-shell electrons, the kineticenergy of which is the difference between the photon energy and 34.5Kev, the K-edge for Xe.

An electron with a given kinetic energy will be able to travel a finiterange in the gas before it collides and ionizes other Xe atoms veryclose to the anode wire. At pressures below 5 atmospheres, the typicaloperating range of an MWPC device, the range of the electron can have avery deleterious effect on the spatial resolution of the image obtainedby detecting the radiation transmitted though an object (i.e., bone) viathe MWPC detector. For example, if a photon interacted at a point in thegas, created a free electron, and the electron traveled to another pointwhere it caused an avalanche at a chamber wire, the initial photoninteraction site in the gas could be one or two wire spacings away fromit initial interaction site. The result is that information as to wherethe initial interaction occurred then becomes very imprecise and, hence,lost of spatial resolution in the image of the object occurs.

It is the accurate identification, as close as possible, of the initialinteraction site between the photon and Xe atom which results in thespatial resolution of the imaged object. Accordingly, it is preferableto minimize the electron range since this range is always added to theintrinsic spatial resolution of the area detector (MWPC), itself. Inpractice, the later can range anywhere from 0.5 to 2.0 mm, depending onchamber design.

It has been discovered that a composite K-edge filter using 2 or more,preferably 3, pure metals stacked on one another, is particularly usefulin modifying the raw spectrum of a radiation source to achieve greaterspatial resolution in the detector. It has been found that metals withatomic numbers from 45 to 52 (usually two metals) combined with at leastone metal whose atomic number ranges from 57 to 65 give an optimumspatial resolution in Xe gas detectors. For example, cadmium (Cd),silver (Ag), and gadolinium (Gd) can be particularly useful incombination. Similarly, cadmium (Cd), tin (Sn) and samarium (Sm) formanother useful combination filter.

The transmission efficiency of a Cd-Ag-Gd filter combination isillustrated in FIG. 9. As can be seen, the filter combination providesfor highest transmissivity in the range of about 42 to about 48 KeV. Theresults of using this Cd-Ag-Gd filter on the raw spectrum of the X-raytube is shown in FIG. 10 where a distinctive narrowing of the energyband can be seen. Similar results in terms of limiting the upper rangeof photons can be achieved for higher X-ray source voltages than 55 Kvby modifying the thickness of the filter elements.

Additionally, the characteristic response of the detector can also beused to further narrow the energy band. In FIG. 11, the efficiency of anMWPC device is plotted, showing that the Xe gas chamber is mosteffective in detecting photons having energies above 35 KeV and below 50KeV. When the filtered X-ray spectra is detected in the MWPC device, theoverall response, as illustrated in FIG. 12, is a well-defined energyband with sharp edges.

The ultimate result of using the composite filter is threefold: first,by blocking low energy photons, it reduces the radiation dose otherwiseabsorbed by the patient and, second, by narrowing the energy band of thephotons from the source, it improves the precision of measurement of theX-ray absorption coefficient for the tissue and bone and, thrid,dramatically diminishes the projected electron range in the detector,itself, thereby improving spatial resolution. The particular advantagesof the composite filter can be seen in Table 1 below, where the X-raytube was controlled so as to impose an upper limit on the photon energyspectrum of 55 KeV and various filter combinations were compared underthe same conditions (i.e., same total metal thickness regardless ofcomposition):

                  TABLE 1                                                         ______________________________________                                        FILTER EFFECTS ON ELECTRON RANGE IN XENON                                                  Filter 1 Filter 2  Filter 3                                      No Filter    (Ag)     (Ag + Cd) (Ag + Cd + Gd)                                ______________________________________                                        electron                                                                              0.92 mm  0.56 mm  0.47 mm 0.36 mm                                     range                                                                         ______________________________________                                    

Another result of the composite filter lies in its ability to reduce theerror typically inherent in bone mineral content (BMC) measurements dueto the presence of fat in the soft tissues of the wrist and forearm. Asdescribed in more detail below, to make bone mineral contentmeasurements, one typically measures the intensity of photons through atissue equivalent bolus which is the same thickness as patient's limband which contains a material having an absorption characteristic whichis the same as tissue free of fat. In actual operations, the fat contentcan be significant (e.g., five to fifteen percent) in heavy individuals;and when it varies, it will lead to errors in measurement from year toyear. Since the composite filter removes a large percentage of thephotons in the range 20 to 35 Kev where the difference between theabsorption coefficients for fat and tissue is largest, the observeddifferences in photon intensity through fat-containing tissue will bereduced substantially for the filtered X-ray spectra.

FIGS. 13-15 illustrate a limb positioning and calibration apparatus 18according to the present invention. In FIG. 13, a top view of theapparatus 18 is shown, including end blocks 60A, 60B, a plurality of topposts 66, a plurality of tracks 78, bottom plate 74 (which is composedof a tissue equivalent material), spring-loaded biasing elements 68A,68B and two side blocks 70, 72, which engage the sides of the patient'slimb in operation. In a preferred embodiment, the side blocks arecomposed of two different materials; for example, side block 70 can becomposed of a bone equivalent material and side block 72 can be composedof a tissue equivalent material. Such materials are well known in thestate-of-the-art of bone densitometry.

FIGS. 14 and 15 show further features of the limb positioning andcalibration apparatus 18 and illustrate its operation. In FIG. 14, theapparatus is shown in its opened position affixed to the housing 20 anddisposed below the radiation source 12. As shown, the apparatus furtherincludes a top cover 62 (at least a portion of which is composed of atissue equivalent material 64) which is attached to the end blocks 60A,60B via a plurality of post clamps 76. In FIG. 15 the apparatus 18 isshown in its closed position.

In operation the limb 16 is positioned in the apparatus 18 such thatside blocks 70 and 72 are pressed by the biasing elements 68A, 68Bagainst the sides of the limb 16 and preferably conform to its shape.For example, the blocks 70, 72 can have a minimum gap, no tension, ofabout 3 centimeters when in the closed position. In operation, theblocks 70, 72 are separated and the limb (e.g. a wrist) is placed in thegap and then the blocks are allowed to close against the limb due toforce of the biasing elements 68A, 68B. To achieve conformation shaping,the blocks can be formed from a flexible plastic shell and filled withfluid, both the plastic and fluid being chosen to provide a calibrationequivalence (e.g. bone equivalence in the case of block 70 and tissueequivalence in the case block 72).

For example, when a wrist measurement is made, the patient's elbow ispositioned so that the wrist and forearm are in approximately a straightline. The liquid-filled flexible material contained in blocks 70 and 72will then form a tight junction and eliminate the gaps near the edge ofthe wrist that otherwise would occur. The bottom part of the wrist willlie on bottom plate 74, which is also preferably a bed oftissue-equivalent material encased in a plastic covering. Next, a topplate 62 which includes an upper tissue equivalence material is loweredonto the wrist 16. The top and bottom plates 64, 74 of the apparatus 18are made flush with the top and bottom of the wrist to give the samethickness as the wrist for the side blocks 70, 72. This is accomplishedby forcing the top plate 62 down to compress the flexible side blocks oneach side of the wrist. The vertical force necessary to compress theside block via the coverplate can be that applied by the operator or thecoverplate itself can also be spring-loaded (not shown). A thin layer offlexible filled material 64 may also be incorporated in the bottomsurface of the top plate 62 so as to ensure a precise fit. Then thecoverplate can be secured by simply turning the screws 76 that lock thecoverplate in place via the posts 66.

As noted above, the side blocks 70, 72 preferably incorporate twoseparate equivalence materials, one for tissue and one for bone, topermit calibration of the system. Such calibration is important toverify results from one machine to another or from measurement sessionto another, perhaps a year or more later. By use of the two equivalencematerials during radiographic imaging, the system can be calibrated tonegate possible changes in the energy spectrum of the radiation sourceor in the detector efficiency.

In one embodiment, calibration can be achieved by taking the naturallogarithm of the intensity of radiation transmitted through the twoequivalence materials. The ratio of these values can then be used toestablish a measurement constant as follows:

    ln (I.sub.b /I.sub.t)=U.sub.b /U.sub.t =k                  (1)

where I_(b) is the intensity of radiation transmitted through a first(e.g., bone equivalence) material, I_(t) is the intensity of radiationtransmitted through a second (e.g., tissue equivalence) material, U_(b)is the bone absorption coefficient, U_(t) is the tissue absorptioncoefficient and k is the measurement constant at which the systemdesirable operates. (It should be clear that the two calibrationmaterials need not be bone and tissue equivalents: any two materials ofdiffering absorption characteristics can suffice. However, bone andtissue equivalency are preferred to assist in imaging resolution andcontrast adjustments, as well).

The calibration apparatus of the present invention thus permits theoperator to measure the ratio of U_(b) /U_(t) and then dynamicallyadjust the system prior to patient imaging until a desired measurementconstant is achieved. Such adjustments will typically involve varyingthe intensity of radiation source (e.g. by adjusting the voltage appliedto the X-ray tube) or modifying the sensitivity of the detector module.Alternatively, any deviation from the desired measurement constant caninstead be noted and used subsequently by the data acquisition or dataprocessing modules to apply a correction factor to the data values. (Itshould also be apparent that the calibration system of the presentinvention is useful not only in single energy photo systems but also inso-called "dual photon" systems which detect radiation at two differentradiation levels to image bones or conduct bone mineral contentanalyses.)

The calibration system of the present invention also permits the user toaccurately determine the thickness of the limb without resort tophysical measurements (e.g., with a caliber). By comparing thelogarithmic intensity values of radiation transmitted through the twocalibration materials, the thickness of the materials can be calculated.Since the clamping means 18 will constraint the materials to have thesame thickness as the limb, the limb thickness can also be inferred.Thus, this thickness T can be calculated according to the followingequation: ##EQU1## where ln (I_(b)) is the natural logarithm of themeasurement radiation intensity transmitted through a first calibrationmaterial (e.g., a bone equivalence medium) and ln (I_(t)) is the naturallogarithm of the measurement radiation intensity transmitted through asecond calibration material (e.g., a tissue equivalent material andU_(b) and U_(t) are the radiation absorption coefficients of the firstand second calibration materials, respectively.

FIG. 16 is a more detailed block diagram of the data readout circuitry100 shown schematically in FIG. 6. The data readout circuitry 100includes gated event digitizers 102A and 102B, highest priority encoders104A and 104B, lowest priority encoders 106A and 106B and adders 108Aand 108B. Each of the X-cathode signal lines are connected to gatedevent digitizer 102A which produces a synchronized set of digitalsignals representative of the X-cathode wires which have "sensed" aradiation event. The highest priority encoder 104A selects the highestX-cathode wire for which a positive digital signal has been generatedand produces a digital number corresponding to the highest wire.Likewise, the lowest priority encoder 106A selects the lowest X-cathodewire for which a positive digital signal has been generated and producesa similar digital number corresponding to the lowest wire. By summingthe outputs of encoders 104A and 106A, and dividing by two, the originof the radiation event on the X-coordinate axis can be determined.

A similar operation is performed by digitizer 102B, encoders 104B and106B, and adder 108B to arrive at the origin of the radiation event onthe Y-coordinate axis. (In practice, one does not need to divide theoutputs of adders 108A and 108B by two, since the sum values will definea new spatial matrix twice the size of the original grid, and thereby,permit recordal of events with greater spatial resolution. The practicalresult is a resolution equal to one-half of the line spacing.)

In FIG. 17, the components of one of the gated event digitizers 102A areillustrated in more detail. As shown, digitizer 102A includes aplurality of amplitude discriminator circuits 120A-120E. (Although fivediscriminator circuits are shown for purposes of illustration, the totalnumber of discriminators in practice will equal the total number ofX-cathode lines.) The lines which feed the discriminator circuits willbe impressed with analog electrical signals of varying amplitudesdepending upon their distance from the radiation event. Thediscriminator circuits are set by an adjustable threshold voltage toproduce a digital output signal when the associated cathode line"senses", an electrical signal above the threshold. The digital outputsof the discriminator circuits 120A-E are registered by latch 124, whichin turn, is triggered by a digital enablement signal generated by asimilar discriminator circuit 122 connected to the anode grid. (Althougha single anode input signal is shown, it should be clear that otherembodiments can include a plurality of anode sensors to detect anegative signal anywhere on the anode grid.) The coincidence of theanode and cathode signals above their respective thresholds yields a setof gated outputs to the encoders 104A and 106A of FIG. 16, therebypermitting a determination of the X-axis location. An identical circuit,with a different number of input and output lines when the detector areais not square, is employed in gated event digitizer 102B of FIG. 16 toobtain the Y-cathode signals and, ultimately the Y-coordinate locationof the radiation event.

The "center of cluster" principle for determining radiation eventlocations in the priority encoders is shown in more detail in FIG. 18.Pulses of varying amplitudes are sensed by the N x-cathode lines. Todetermine the center of the signal cluster (X) and, hence, the origin onthe X-axis of the radiation event, the priority encoders first determine(X1) and (X2), the locations of the end points of the pulsedistribution. The center (X) can be calculated as follows: ##EQU2##

Again, an analogous calculation is made for the M Y-cathode lines todetermine the center of the Y-cathode signals.

FIG. 19 is a more detailed block diagram of the data acquisition module22 shown schematically in FIG. 2. The data acquisition module 22includes programmable logic device 130, address multiplexers 132A, 132B,incrementer 134, data multiplexers 136A, 136B, memory and multiplexercontrollers 138A, 138B, dynamic random access memory banks 140A, 140B,logic controller 142 and a computer interface controller 144.

As shown in FIG. 19, the data acquisition module is formed from twoseparate memory subassemblies, each consisting, for example, of astandard 512×512×16 bit memory operated by a common programmed logicdevice 130 and logic controller 142. When a given memory is filled, thelogic controller 142 directs new events to the other memory bank andproceeds to transfer the contents of the first memory bank through thecomputer interface controller 144 to a data processing module, a datadisplay module or a data storage module.

The memory banks 140A, 140B each can each be constructed as a single512×512 pixel image memory, or they can be partitioned to representmultiple smaller images, such as for example, four 256×256 pixel imagesor sixteen 128×128 images. Any of these arrangements permits the systemto take multiple frames of data sequentially in time; each memory bank(or partition) thus acts as a histogram creating an image based uponradiation events recorded over time. Since the module 22 permits onememory bank to be transferred as the other is filled, it is possible toprocess or store data continuously.

The interface controller 144 is preferably a small computer systemsinterface (SCSI) which is compatible with a wide variety of computerhardware, thereby allowing data transfer continuously, in either frameor list mode, to an external computer memory and/or an external magneticor optical memory hard disk. Moreover, since the SCSI standard supportsmultiple SCSI devices on a single bus, it is possible to have severaldata acquisition systems operating in parallel from the same or multiplesources.

The logic controller 142, incrementer 134 and programmed logic device130 of FIG. 19 cooperate to acquire and transfer data from the detectormodule, which can be a multiwire proportional chamber device, asdescribed above, or another type of imaging device, such as a gammacamera or other photon imager. The logic controller 142 can beconstructed from a general purpose microprocessor, such as the HD64180microprocessor manufactured by Hitachi (San Jose, Calif.) and/or othercommercially available parts. The microprocessor is preferablyconfigured to define a direct memory data access channel for fasttransfer of data between the memory banks 140A, 140B and the computerinterface controller 144, and serves to refresh the dynamic randomaccess memories periodically as well as initialize the programmed logicdevice 130, and multiplexers 132 and 136 and activate the interfacecontroller 144.

The microprocessor is programmed to perform tasks in response torequests by either the computer interface controller or other eventsthat signal the end of a data acquisition cycle. In a typical sequence,the computer interface controller 144 will alert the logic controller142 of a pending host computer request. This request can be, forexample, an order to commence acquiring data from a wire chamberdetector. The logic controller 142 will then decide which of the memorybanks is to be used for this data acquisition cycle and set up thememory controllers 138 for that purpose. The logic controller 142 willalso signal the programmed logic device 130 to begin acquiring data asdescribed below. In addition, the logic controller 142 can initiate aninternal timer to stop the data acquisition process after a predefinedtime has elapsed. Alternatively, it can program an event counter to stopdata acquisition after a predefined number of events have been recorded.

A host computer may also request from the logic controller 142 theinitiation of data transfer from a previously acquired frame. The logiccontroller 142 will then direct the memory controllers 138 andmultiplexers 136 and 132 to connect the corresponding memory bank 132Aor 132B to the data path of the logic interface controller 144. Thelogic interface controller 142 will then set up the parameters of adirect memory access controller to sequentially read the locations ofthe memory bank and simultaneously write them to the computer interfacecontroller 144.

The incrementer 134 can be implemented by a parallel load counter, suchas the model 74F779 counter manufactured by National Semiconductor,Fairchild Division (Mountainview, Calif.). The programmed logic device130 can also be implemented with a commercially available, programmablelogic device such as the model 20R4 device manufactured by AdvancedMicro Devices (Sunnyvale, Calif.).

The various multiplexers and memory elements also shown in FIG. 19 areconventional in design and can be implemented with a wide variety ofcommercially available components. It should be clear that variouschanges, additions and subtractions to the circuitry of FIG. 19 can bemade by those skilled in the art without departing from the spirit orscope of the invention. For example, the memory and mutiplexer controlfunctions of elements 138A, 138B can be supported by separate bit sliceprocessor elements or can be implemented by firmware as part of thelogic controller 142 or the programmed logic device 130. Obviously, thefunctions of the components illustrated in FIG. 19 can be combined intolarger integrated circuits or divided into smaller processing elements.

In one preferred embodiment, the programmed logic device 130 isimplemented as a state machine with eight states. The device 130 permitsvery fast acquisition of histogramming information by addressing alocation in one of the memory banks, which corresponds to the particularcoordinate location in the detector module where the radiation eventoccurred, retrieving the contents of the memory location and presentingit to the incrementer 134 which adds a "1" to those contents, and thenrewrites the new number to the addressed memory contents.

The functions of programmed logic device 130 can be further understoodby reference to the timing diagrams of FIG. 20A and further illustratedby the schematic diagram of FIG. 20B. As shown, the state machine hasthe following sequence:

In the idle mode, the state machine cycles between state #0 and state#1. If the detector module signals a new event by asserting the INFOREADY line during state #0, the state machine switches to the activemode.

State #1 is a timing safeguard to insure that the INFO READY signal hasbeen received.

In state #2, the horizontal (or x) information is requested from thedetector module, which is latched as row information in the DRAM-basedmemory bank 140.

In state #3, the vertical (or y) information is requested from thedetector module, which is latched as column information in the memorybank 140. At this point, the present contents at the location X-Y in thememory bank 140 are retrieved. The output buffers are enabled, and theincrementer is ordered to load the contents (READ MEMORY signal).

State #4 is a timing delay to ensure that the memory contents have beenloaded into the incrementer.

By state #5, the present contents of the X-Y position have been loadedin the incrementer 134, and an order to INCREMENT has been given.

State #6 is again a timing delay to insure that the incrementer hascompleted it task.

In state #7, an order is given to the incrementer 134 to output the newdata, while simultaneously writing said data to the position X-Y in thememory bank 140 (MEMORY LOAD signal).

The programmed logic device can be driven at any clock rate compatiblewith the other components of the system. For example, when a 10 MegaHzclock is employed, the state machine can cycle through its eight statesas shown in FIG. 20A in four clock cycles, or approximately 50nanoseconds per state. A faster clock can also be employed to furtherreduce the data acquisition time.

The input signal labeled REFADV indicates to the state machine that aDRAM-refresh cycle is about to take place. The state machine refrainsfrom initiating the 8 state sequence while this signal is asserted.

The DATA READY signal, used by the wire chamber system to indicateavailability of data, may be withdrawn by state #3, and must bewithdrawn by state #7.

While reading the histogramming memory 140, the general purpose logiccontroller 142, and not the programmed logic device 130, preferably,generates the necessary control signals to retrieve the information fromthe memory 140 and presents it to the processor interface 142. When datatransfer from a memory element to the computer interface 144 is takingplace, no data acquisition can be performed to that histogram memory andvice versa.

With reference again to FIG. 19, the data acquisition module 22 providesfor simultaneous buildup of a histogram, and transfer of data to theinterface 144 of a previously-acquired histogram, by employing twoidentical histogramming memory banks 140A, 140B. This prevents loss ofdata acquisition information while reading previously acquiredinformation.

A plurality of two-position multiplexers or digital switches areincorporated for this purpose. They provide a connection of each of thehistogramming memory banks 140A, 140B to the data acquisition or datatransfer address and data lines. Two unidirectional multiplexers 132A,132B--labeled ADDRESS MUX--route either the wire chamber information orthe logic controller generated address to the corresponding memorybanks. Two bidirectional multiplexers 136A, 136B--labeled DATAMUX--route the histogram information to the incrementer 134 (when in adata acquisition cycle) or to the processor interface 144 (when in adata transfer cycle).

The multiplexers, for one of the histogramming memory banks, are alwaysin the opposite position, as the ones for the remaining histogrammemory. The one exception is during the DRAM refresh cycle (generated bythe logic controller 142), in which both address multiplexers 132A, 132Bare connected to the logic controller 142.

FIG. 21 illustrates the data flow path during a typical portion of theacquisition cycle. In this figure, the acquisition module 22 is shown ina state where memory bank 140A is acquiring data, while memory bank 140Bis transferring data through the processor interface 142.

Returning to FIG. 2, the remaining elements of system 10, the dataprocessing module 24, a data display module 26 and a data storage module28 can all be assembled from commercially available components. Forexample, a wide variety of general purpose microcomputers, minicomputersor mainframe computers can be employed to process the data in the dataprocessing module 24. Similarly, a wide variety of computer monitors orhigher resolution video display equipment can be employed in the displaymodule 26. In accordance with conventional image generation techniques,the image presented to the operator can be composed of pixels havinggray-tone levels or color values which are representative of thedetector responses at given areas of the patient's limb being imaged.The data storage module 28 can also be of conventional design and relyon any one of a number of known storage media, such as magnetic disks,diskettes, tape cassettes, optical disks, or solid-state random accessmemory elements.

In one preferred embodiment, the data processing module 24 and thedisplay module 26 can be used cooperatively to perform measurements ofbone density on a particular bone structure or bone structures of thepatient's limb being imaged. A procedure for such imaging and bonedensity measurements, which can be implemented by software, firmware, ora combination thereof, in the data processing module 24, follows withreference to FIGS. 22-24, which are illustrative of the bone structuresof a patient's wrist.

In FIG. 22, an overall schematic view of the presentation of a patient'slimb 16, in this instance, a wrist, to the system is shown. The bonestructure of the patient's limb 16 in shown in phantom, except for thatportion 150 which is actually viewed by the detector. As can be seenfrom this schematic illustration, the two principal bone structures inthe wrist image are the radius 152 and the ulna 154. The data processingmodule of the present invention employs the styloid tips of the radiusand the ulna bones as reference points, and establishes measurementareas in select regions of those bones that contain high percentages oftrabecular bone, rather than cortical bone. These select regions, in thecase of the radius and the ulna, have been determined to lie at specificdistances from the respective styloid tips of these bones.

In practice, to improve sensitivity in detecting bone loss, it isimportant to measure areas that contain high percentages of trabecularbone. Thus, in accordance with the invention, measurement areas areselected that contain 60% or more of trabecular bone. The percentage oftrabecular bone present in the radius and the ulna beyond 30 millimetersfrom the styloid tip of each bone drops below 10% and the percentage ofcortical bone for both the radius and the ulna is approximately 95% atpoints beyond 40 millimeters from the styloid tip.

It has also been determined that the region associated with this 60%limit of trabecular bone usually begins at 2 millimeters and 4millimeters, respectively, from the styloid tips of the radius and theulna, respectively. The highest trabecular content by volume istypically found about 9 millimeters to 20 millimeters from the styloidtip in the radius and between about 4 millimeters and 12 millimetersfrom the styloid tip in the ulna. In accordance with the invention,therefore, bone mineral content measurement regions are selected withineach of these areas of high trabecular bone content.

In use, the operator can perform a bone mineral content measurement byemploying the data processing module 24 and data display module 26together. In one typical protocol, the system is initialized and animage is displayed to the clinician via a conventional computer monitor.The resolution of the display is not critical, so long as the bone edgesare identifiable by the operator. In one preferred embodiment, theoperator is prompted to use a cursor control device, such as aconventional mouse or joy stick, to place the styloid tip region markersaround the respective styloid tips of the radius and the ulna. Thesemarkers may include, for example, hollow rectangular region markers,similar to styloid tip markers 166 and 168 shown in FIG. 23, whichdepicts a more detailed radiological image of a human wrist. In theillustrated embodiment, the data processing module executing the methodsteps of this invention reads the position of the styloid tip regionmarkers to generate initializing data for the location of the styloidtips.

The location of edges surrounding the styloid tips within each markerregion can be determined by techniques known in the art. Bone edgefinding can be implemented in accordance with any conventional edgefinding algorithms, such as those disclosed in J. F. Canny, Tech. Rep.720. Artificial Intelligence Laboratory, MIT (1983); "A ComputationalApproach To Edge Detection," Vol. PAMI-8 IEEE Trans. Pattern Anal.Machine Intell. (1986) or F. Bergholm, "Edge Focusing," Vol. PAMI-9 IEEETrans. Pattern Anal. Machine Intell. (1987). Preferred edge findingalgorithms include those which reduce noise contributions in the imagesand optimize the location of single edges rather than multiple spuriousedges.

The edge finding step is followed by a positional search executed on theedges of the styloid tip to identify one or more pixels that define theextremity of the styloid tip on each bone. Because the precisepoint-by-point shape of the respective radial and ulnal styloid tips candiffer from individual to individual, it can be useful to employ morethan one form of positional search or a curve fitting routine todetermine the extremity of each styloid tip. The search or curvedfitting routine can include a technique which finds the minimum-gradienttangent between points defining the edge, or one which locates theintersection of curves fitted to the styloid tip edges.

As an alternative to reading and responding to operator input andfinding the styloid tips, the invention can utilize known techniques forinitializing edge finding on the outer edge of the radius or ulna andprogressing up and around the styloid tips to determine the edges of thebone extremities. However, utilization of operator input reducesprocessing time. Moreover, the step of responding to operator inputincreases accuracy over a fully automated system, because it is likelythat bone edges other than those of the radius and ulna will be presentwithin the styloid tip region of interest. The phenomenon is illustratedin FIG. 23, where the edges of the scaphoid bone 156 near the radius,and the lunate bone 158 near the ulna, are proximate to the styloid tips162 and 170. When the styloid tip locations have been determined, one ormore pixels identifying the extremity of the tip can be highlighted onthe display for inspection by the operator, to confirm properidentification of the styloid tips.

Next, the data processing module executes a further edge finding step todetermine the inner and outer longitudinal edges of the radius and theulna. (As used herein, the term "longitudinal direction" is meant todefine the direction corresponding to the long axis of the bone. Forpurposes of illustration, this longitudinal axis corresponds to theY-coordinate axis of the detector.)

The initial starting position can be, for example, 3 centimeters fromthe styloid tip extremity of the radius. This initial position ispreferable to ensure that the edge finding occurs in a region having aninterosseous gap between the radius and the ulna. Typically, at thisposition with the wrist in a neutral position (i.e., with no twisting ofthe wrist off the X-reference plane) the interosseous separation betweenthe ulna and the radius can range from 0.5-2.5 centimeters, dependingupon the size and shape of the bones. Thus, by initiating the edgefinding process at a selected displacement, for example, approximately 3centimeters, from the extremity of the styloid tip of the radius,ensures that the edge finding will occur in a region having a gapbetween the radius and the ulna.

Longitudinal edge finding thus progresses along the bone from theinitial position for a selected distance in the direction of a styloidtip and from the initial position for a selected distance in theopposite direction, to identify the edges of the bones on either side ofthe initial starting position the selected distance can be, for example,2 centimeters or less from this initial position.

When the longitudinal edge finding process is complete, the widths ofeach bone can be computed at selected Y-axis positions. A horizontalline, referred to as the "region base line," is then established anddisplayed at a Y-axis position on each bone having a selecteddisplacement from the extremity of the respective styloid tip. A bonemineral content measurement can then be made on each bone. Themeasurement regions can be highlighted on display for the inspection ofthe operator.

As illustrated in FIG. 24 the measurement regions 170 and 172 can be,for example, rectangular in form having a selected height and a selectedwidth based on the bone widths calculated previously. In particular, themeasurement regions can extend from bone edge to bone edge, as doesmeasurement region 172 on the ulna 154 in FIG. 24 or, the measurementregions can be reduced by a selected amount, to eliminate the outer andinner edges of the imaged bone from the BMC measurement region. Area 170on the radius 152 in FIG. 24 is an example of a narrowed measurementregion.

In many cases, narrowing the measurement region provides increasedsensitivity in BMC measurements by reducing inaccuracies in measurementand computation which can result from highly irregular bone shapes.

The BMC values can be derived in a known manner, by summing pixelintensity values in each measurement region. Bone density values canalso be obtained by dividing the BMC value by the size of themeasurement regions. The area covered by the measurement regions isproportional to the number of pixels in the measurement region. Thoseskilled in the art will appreciate that BMC values are proportional tothe sum of the values of the data points in the measurement region,which in turn are represented by the pixel intensity values in themeasurement region. In particular, the digital number associated witheach pixel is typically proportional to the magnitude of the radiationreceived by the detector in the region corresponding to that pixel.

Processing can be terminated when the BMC and density values have beenobtained. Alternatively, the process can be re-initiated for processingof additional radiographic images.

Additional techniques for imaging bone structure and for deriving bonemass in density can be found in commonly owned U.S. patent applicationSer. No. 321764 entitled "Methods And Apparatus For Bone Measurement"filed on even date herewith and incorporated herein by reference.

The present invention also permits a number of additional analyses to beperformed from the stored data. For example, in successive bone mineralmeasurements of the same patient, the currently generated image data canbe compared with previously stored images. A comparison can be madebetween the current orientation of the radius and ulna and theorientation of the radius and ulna represented in the stored data.

In a further preferred embodiment of the invention, a correction factoris established for the X and Y coordinate factors, to correlatepreviously generated and currently generated images and to correct smallerrors in the positioning of the patient's limb. Two-dimensional images,such as the bone joint images processed in accordance with the presentinvention, have defined orientations within an X-Y orthogonal coordinateaxis system. Differences in the orientation of the ulna and the radiusin successive images can be detected with reference to this X-Ycoordinate system.

One technique is to evaluate the gradient at various points along theouter edges of each bone, to identify the flattest portion of the outeredge of each bone. Curve-fitted portions of the currently generated andpreviously generated bone joint images can than be compared to determinewhether small translations and/or rotations of the imaged bones haveoccurred from one measurement to the next.

While translational transformations in the direction of the Y-axis donot significantly affect the repeatable selection of measurement regionsin successive images, rotational transformation should be corrected fromimage to image. Corrections for rotational transformations can be madeby computing the angles between the flat portion of the outer bone edgesfor each of the first and second images. Based on the computed angles,corrections for both the X- and Y-axis positional pixel values can beprovided by known computational techniques. Evaluation of these angles,and the positional values of the extremities of the styloid tips,permits images to be superimposed, or otherwise placed intopoint-to-point registration, to ensure that the vertical and horizontalbone distances are consistent from image to image for a givenindividual.

It will be understood that changes may be made in the above constructionand in the foregoing sequences of operations without departing from thespirit or scope of the invention. It is accordingly intended that allmatter contained in the above description or shown in the accompanyingdrawings be interpreted as illustrative rather than in a limiting sense.

Having described the invention, what is claimed as new and secured byLetters Patent is:
 1. An apparatus for analyzing biological structuresby photon absorptiometry, the apparatus comprising:a radiation sourceemitting photons with a range of energy levels, said radiation sourceincluding at least one source filter element which operates to reducethe emission of high energy photons; and a detector means fordetermining the spatial intensity of radiation from the radiationsource, said detector means having a selective characteristic responsewhich is relatively insensitive to low energy photons, the detectormeans cooperating with said source filter to measure the intensity ofphotons within a narrowed band of energy levels.
 2. The apparatus ofclaim 1 wherein the radiation source is an X-ray source.
 3. Theapparatus of claim 1 wherein the radiation source is a broad band X-raysource emitting X-rays at energy levels up to about 70 KeV.
 4. Theapparatus of claim 1 wherein the source filter comprises at least oneselectively transmissive filter element which reduces the transmissionof photons having energy levels above about 50 KeV.
 5. The apparatus ofclaim 1 wherein the source filter element is a selectively transmissivefilter element comprising lanthanide metals and alloys thereof.
 6. Theapparatus of claim 1 wherein the source filter element further comprisesa gadolinium sheet element.
 7. The apparatus of claim 6 wherein thegadolinium sheet element has a thickness ranging from about 0.01 toabout 1.0 millimeters.
 8. The apparatus of claim 1 wherein the sourcefilter element further comprises a samarium sheet element.
 9. Theapparatus of claim 8 wherein the samarium sheet element has a thicknessranging from about 0.01 to about 1.0 millimeters.
 10. The apparatus ofclaim 1 wherein the detector means has a characteristic responsive whichis relatively insensitive to photons having an energy level below about35 KeV.
 11. The apparatus of claim 1 wherein the radiation sourcefurther comprises at least one additional filter element which operatesto reduce the emission of low energy photons.
 12. The apparatus of claim11 wherein the second filter element is a selectively transmissivefilter element which reduces the transmission of photons having energylevels below about 35 KeV.
 13. The apparatus of claim 11 wherein theadditional filter further comprises a selectively transmissive filterelement selecting from the group consisting of rhodium, silver, cadmium,indium, tin, tellurium and alloys thereof.
 14. The apparatus of claim 13wherein the additional filter element is a metal sheet element having athickness ranging from about 0.01 to about 1.0 millimeters.
 15. Theapparatus of claim 11 wherein the additional filter element is acomposite filter element comprising a plurality of metal sheets selectedfrom the group consisting of rhodium, silver, cadmium, indium, tin,tellurium and alloys thereof.
 16. The apparatus of claim 15 wherein theadditional plurality of metal sheets each have a thickness ranging fromabout 0.01 to about 1.0 millimeters.
 17. The apparatus of claim 15wherein the composite filter element further comprises sheets of silverand cadmium.
 18. The apparatus of claim 15 wherein the composite filterelement further comprises sheets of tin and cadmium.
 19. The apparatusof claim 1 wherein the detector means further comprises a multiple wireproportional counter filled with an ionizable gas.
 20. The apparatus ofclaim 19 wherein the ionizable gas further includes xenon gas.
 21. Theapparatus of claim 1 wherein the apparatus further comprises acollimating means for collimating the photons from the radiation sourceafter transmission through the source filter.
 22. An apparatus foranalyzing biological structures by photon absorptiometry, the apparatuscomprising:a radiation source emitting photons with a range of energylevels less than about 60 KeV; a first source filter comprising at leastone selectively transmissive material which operates to reduce theemission of high energy photons from the radiation source above about 50KeV; a second source filter comprising at least one selectivelytransmissive material which operates to reduce the emission of highenergy photons from the radiation source below about 35 KeV; acollimating means for collimating the photons from the radiation sourceafter transmission through the first and second source filters; and amultiple wire proportion counter detector means for determining thespatial intensity of radiation from the radiation source, said detectionmeans including ionizable xenon gas having a selective characteristicresponse to ionizable radiation which is relatively insensitive to lowenergy photons, the detector means cooperating with said source filtersto measure the intensity of photons within a narrowed band of energylevels.